Method for detecting biomolecules electrically and biochip therefor

ABSTRACT

The present invention relates to a method for detecting the presence and/or the reaction of a biomolecule by monitoring changes of electrical property accurately according to the biological, biochemical or chemical reaction of the biomolecule, and a biochip provided for this purpose.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. patent application Ser. No.12/922,934, filed on Sep. 16, 2010, which is a 371 of PCT/KR2009/000285filed on Jan. 20, 2009, which claims the benefit of Korean ApplicationNo. 10-2008-0027099 filed on Mar. 24, 2008, the contents of each ofwhich are incorporated herein by reference.

TECHNICAL FIELD

The present invention relates to a method for detecting the presenceand/or the reaction of a biomolecule by monitoring changes of electricalproperty accurately, after the biological, biochemical or chemicalreaction of the biomolecule and a biochip provided therefor.

Particularly, the present invention relates to a method for detectingchanges of electrical property, especially the changes of an impedanceor capacitance value in a reaction chamber sensitively after thebiological, biochemical or chemical reaction of the biomolecule, byusing a sensing electrode installed in the reaction chamber and abiochip provided for this purpose.

More particularly, the present invention relates to a method fordetecting changes of electrical property such as changes of an impedanceor capacitance value sensitively, after a receptor immobilized onto asensing electrode of a reaction chamber reacts with a biomolecule as atargeting molecule and a biochip used for this purpose.

BACKGROUND ART

Biochip is a device that can analyze genetic information and proteininformation automatically in a large scale, or detect the presence andthe function of a biomolecule easily and rapidly. This biochip is beingactively applied for various fields including gene and proteinresearches, medicines, and agricultural, environmental and chemicalindustries, etc.

The biochip is classified broadly to genotyping chip, expression chipand microfluidics chip: the genotyping DNA chip is to detect thepresence of a particular gene by using a probe; the expression DNA chipis to monitor the expression profiling of gene associated with aparticular disease; and the microfluidics chip is to detect the presenceand/or the reaction of a biomolecule within a sample including blood andurine. Presently, the genotyping DNA chip is commercialized and usedwidely in research areas and medical diagnosis areas.

Generally, the term microarray chip defines a chip that arrays hundredsto ten thousands kinds of genes or proteins mounting on a glass plate byusing a microarray apparatus. Among these, DNA chip is to microarrayoligonucleotides as probes on a glass plate in order to identify thepresence of a particular gene with a fluorescence scanner.

The DNA chip is being utilized practically for research and diagnosisfields. Particularly, this chip is applied to elucidate the genefunction including cellular metabolism, physiological phenomena andmutual relation between genes by using gene expression profiling andgenotyping techniques, etc. The DNA chip is also used widely indiagnosis to examine a mechanism causing a particular disease such ascancer, prognostic diagnosis and action of drugs, to identify geneticinformation of microbes causing diseases, and to screen mutations, etc.

Diagnostic DNA chip has been developed in 1994 by Dr. Steve Fodor inAffymetrix Co. Ltd., and the first HIV gene chip started to becommercialized in a market. Nowadays, researches upon the diagnostic DNAchip are attempted actively in order to diagnose chronic diseasesincluding HIV, rheumatism, autoimmune disease, chronic nephritis,atherosclerosis, atopic dermatitis and allergy, etc. Especially, recentstudies upon DNA chips tend to develop chips for diagnostic use ratherthan chips for research use. Moreover contrasting to genotyping chipsuseful for diagnosing genotype, pathogen and virus, an approach on geneexpression profiling capable of diagnosing various diseases includingcancer and leukemia, is being accomplished.

Recently, microfluidics chip (Lab-on-a chip) that can detect a lot ofdiseases coincidently from one trial and predict outbreak of diseasesfrom genetic information of an individual by introducing IT and nanotechnologies, attracts attention. The microfluidics chip is alsoreferred to as biochip. This chip is used to analyze a reactionprofiling of various biomolecules within a chip, after a minute amountof an analytic target material (DNA, RNA, peptide, protein, etc.) isintroduced into a chip chamber. This biochip is to detect the presenceand/or the reaction of a biomolecule by monitoring changes of electricalproperty from an electrode installed in a chip, after being reacted withthe biomolecule in a reaction chamber.

Such a biochip is highly applicable for medical diagnosis, because itcan identify the presence and/or the reaction of a biomolecule moreeasily and rapidly by detecting electrical signals than any other DNAchips mentioned above.

For example, HPV DNA chip is a device that prepares HPV oligonucleotideprobes and microarrays these probes on a glass plate in order todiagnose whether HPV, a pathogenic virus causing cervical cancer ispositive or not. Nevertheless, it is impossible to directly diagnose thepositive status of HPV, right after suspected sample is collected. Thatis to say, primers for amplifying HPV viral gene, labeled withfluorescence should be prepared in advance. Then, the collected sampleshould be amplified by performing a PCR, mounted onto an HPV DNA chipand monitored to examine a fluorescent signal with a fluorescencescanner. However, this system for detecting hybrids by using a DNA chip(laser-induced fluorescence) is inconvenient to be manipulated andspends time a lot. Therefore, this method has various problems anddisadvantages. It needs high cost due to labeling a DNA sample with afluorescent material and is not portable because of using an expensivefluorescence scanner.

In contrast, the biochip can identify the presence and/or the reactionof a biomolecule relatively easily and rapidly by detecting electricalsignals. Particularly, the biochip can detect the presence and/or thereaction of a biomolecule (for example, DNA) by using electrical signalsrather than fluorescent signals. More particularly, the biochip adopts asystem for detecting changes electrically, in which the change of animpedance value (or a capacitance value) is monitored after reacting areceptor immobilized onto an electrode with a biomolecule, or the changeof an impedance value (or a capacitance value) is monitored afterreacting between biomolecules in a chip chamber.

For example, it is reported in the PCR process that dNTP should degradeto dNMP and diphosphate and the resulting dNMP is polymerizedsimultaneously from a primer complementary to a DNA template sequence soas to synthesize DNA. Accordingly, the impedance value within a PCRreagent increases as DNA concentration increases (See Korean PatentPublication No. 10-2004-0042021). Therefore, it is possible to determinewhether the PCR reaction is performed and a particular DNA sequenceexists or not, when a PCR reaction chamber is manufactured with abiochip structure and the changes of an impedance value in a reagent aredetected electrically on an electrode installed in such a biochip.

In addition, it is possible to detect electrically whether a PCRreaction or a hybridization reaction is accomplished and/or whether atarget nucleotide sequence exists or not, when oligonucleotides such asprimers or probes are immobilized on a biochip electrode, a PCR reactionor a hybridization reaction is conducted and then, the changes of acapacitance or impedance value are measured with the electrode on whichthe oligonucleotides are immobilized.

In general, the biochip adopts a system for determining the reaction andthe presence of a particular biomolecule by monitoring the impedancechange with an electrode equipped in the biochip after reacting thebiomolecule and other reagents (for example, the receptor immobilizedonto an electrode). Accordingly in order to guarantee the reliability ofa biochip, it is important to measure the change of the impedance valueaccurately. Generally, impedance (Z) indicates the sum of resistance (R)as a real number portion and reactance (X) as an imaginary numberportion (See following Mathematical formula 1), and the magnitude ofimpedance corresponds to a square root of resistance score (R) andreactance score (X) (See following Mathematical formula 2).

$\begin{matrix}\begin{matrix}{Z = {R + {j\; X}}} \\{= {R + {j\left( {X_{L} - X_{C}} \right)}}} \\{= {R + {{j\left( {{\omega \; L} - \frac{1}{\omega \; C}} \right)}\mspace{14mu}\lbrack\Omega\rbrack}}}\end{matrix} & \left\lbrack {{Mathematical}\mspace{14mu} {formula}\mspace{14mu} 1} \right\rbrack \\\begin{matrix}{{Z} = \sqrt{R^{2} + X^{2}}} \\{= {\sqrt{R^{2} + \left( {{\omega \; L} - \frac{1}{\omega \; C}} \right)^{2}}\lbrack\Omega\rbrack}}\end{matrix} & \left\lbrack {{Mathematical}\mspace{14mu} {formula}\mspace{14mu} 2} \right\rbrack\end{matrix}$

Accordingly, the impedance (Z) value is made to have a correlation withreactance (X). Also, the reactance (X) has a correlation with thecapacitance (C) value because it is ωL−1/ωC. Therefore, the change ofthe capacitance (C) value varying according to biological, biochemicalor chemical reactions, is reflected by the change of the impedancevalue, which enables to check the reaction and/or the presence of abiomolecule after its measurement. Finally, it is verified that thechange of the capacitance value should change the impedance value in abiochip and influence upon the sensitivity of the biochip.

The change of the impedance value can be affected by two kinds ofcapacitance values after reacting the biomolecule and other reagents(for example, the receptor immobilized onto an electrode). Particularly,when the change of the capacitance value is measured by the electrode ofa biochip, total change of capacitance value (C_(T)) in a biochip caninclude change of a capacitance value (C_(d)) on the surface of asensing electrode after reacting a receptor immobilized onto anelectrode with a biomolecule, and a capacitance value (C_(i)) between animaginary electrode plate and another sensing electrode, which isdifferent from the change of the capacitance value (C_(d)) (Seefollowing Mathematical formula 3).

$\begin{matrix}{C_{T} = \frac{1}{\frac{1}{C_{t}} + \frac{1}{C_{d}}}} & \left\lbrack {{Mathematical}\mspace{14mu} {formula}\mspace{14mu} 3} \right\rbrack\end{matrix}$

In a biochip, the biomolecule is reacted under a reagent filling areaction chamber and the reagent generally contains a buffer solution orelectrolytes. The buffer solution and electrolytes contain ions toinfluence conductance between sensing electrodes of a biochip. Thisbuffer effect upon the conductance between sensing electrodes furtheraffects the capacitance value (C_(t)) between the imaginary electrodeplate and another sensing electrode, which are contained in the bufferfunctioning as a dielectric substance. Finally, this influences totalchange of the capacitance value (C_(T)).

On the other hand, the capacitance value depends on the dielectricconstant of a material between electrodes as defined in followingMathematical formula 4. It is verified that when the dielectric constantof a material between electrodes is small, the capacitance value becomessmall.

$\begin{matrix}{C = {\frac{Q}{V} = {ɛ\; \frac{A}{t}}}} & \left\lbrack {{Mathematical}\mspace{14mu} {formula}\mspace{14mu} 4} \right\rbrack\end{matrix}$

A: area of electrode

T: interval between electrodes

ε: dielectric constant of material between electrodes

However, the reagent such as the above-mentioned buffer has a smalldielectric constant due to the ion conductivity and thus, 1/C_(t), anelement comprising total change of the capacitance value (C_(T)) becomeslarge in Mathematical formula 3. Accordingly, when the biomolecule isreacted under a buffer solution filled in the reaction chamber of abiochip, the capacitance value (C_(d)) changed by the reaction betweenthe sensing electrode and the biomolecule can be absorbed into thecapacitance value (C_(t)) between the imaginary electrode plate and theanother sensing electrode contained in the buffer functioning as adielectric substance. Particularly as demonstrated in the Mathematicalformula 3, the capacitance value (C_(d)) changed on the surface of thesensing electrode by reacting the receptor immobilized onto theelectrode with the biomolecule does not influence the total change ofthe capacitance value (C_(T)), because the reciprocal number (1/C_(t))of the capacitance value (C_(t)) between the imaginary electrode plateand the another sensing electrode filled with the reagent containingbuffer ions as a dielectric substance is large. Therefore, it isdifficult problematically to determine the presence and/or the reactionof a biomolecule by the electrical detection using the sensing electrodeof a biochip.

Referring to this, FIG. 1 illustrates that the curve of the impedancevalue varied according to frequency is observed to have the same trendbefore and after PCR reaction, even though a particular target DNAtemplate as a diagnostic sample exists. If conventional biochip couldmeasure PCR reaction using a particular target DNA template as abiomolecule accurately, the change of the capacitance value (C_(d)) onthe surface of a sensing electrode could be detected electrically afterthe PCR reaction between a primer fixed onto the sensing electrode and atemplate ssDNA. The change pattern of impedance curves might also beobserved before and after the PCR reaction, because the change of thecapacitance value (C_(d)) on the surface of the sensing electrode shouldbe reflected after DNA amplification by the PCR reaction.

However, since the capacitance value (C_(d)) changed after reacting asensing electrode and a biomolecule may be passed over by thecapacitance value (C_(t)) between an imaginary electrode plate andanother sensing electrode as described above, the impedance curve is notchanged and just observed as a single curve before and after the PCRreaction, when being measured with an biochip-connected apparatusmonitoring an impedance value (See FIG. 1). Therefore, in biochips ofprior arts, the capacitance value (C_(d)) changed by reacting a receptorimmobilized onto an electrode and a biomolecule, is passed over by thecapacitance value (C_(t)) between an imaginary electrode plate andanother sensing electrode containing a buffer solution as a dielectricsubstance, and thus, two curves of the changes of the impedance valuewhich mean total change of the capacitance value (C_(T)) are notmeasured. Problematically, it is difficult to monitor whether thebiomolecule reacts and/or exists or not in practice.

Particularly, in US Patent Publication No. 2004/0110277, the system forsensing a hybridization reaction, in which a probe DNA is immobilizedonto an electrode of sensor cell in a biosensor and hybridized with atarget DNA to change a capacitance value and finally, a current value ofa transistor, has been disclosed. However, it does not focus to settlefollowing problem: since the capacitance value (C_(d)) changed byreacting a receptor immobilized onto an electrode with a biomolecule ispassed over by the capacitance value (C_(t)) of the buffer solution, thereaction between the receptor immobilized onto the electrode and thebiomolecule as a target material cannot be detected electrically.

Besides, in US Patent Publication No. 2006/0226030, the technique forsensing a hybridization reaction by detecting a capacitance value with asensor comprising a plate, an electrode and a catcher moleculeimmobilized on the plate, in which the catcher molecule is hybridizedwith DNA single strand labeled with metallic ball (having a differentdielectric constant from that of the other material) and as a result,the label surrounds the electrode to change a part of the capacitancevalue regarding an electrode impedance value, has been disclosed.However, it did not pay attention to overcome following problem: sincethe capacitance value (C_(d)) changed by reacting a receptor immobilizedonto an electrode with a biomolecule is passed over by the capacitancevalue (C_(t)) of a buffer solution, changes of electrical property cannot be detected sensitively during actual reaction.

Furthermore, in conventional techniques, a system for increasing sensingsensitivity of IDE by reducing the interval of sensor electrodes inorder to improve the sensing sensitivity of a biochip, has beendisclosed. But, it did not recognize the problem mentioned above: sincethe capacitance value (C_(d)) changed by actual reaction is passed overby the capacitance value (C_(t)) of a buffer solution, the changes ofthe capacitance value and the impedance value cannot be detectedsensitively.

Therefore, conventional biochips and methods using the same foridentifying the reaction and/or the presence of biomolecules disclosedin prior arts should be improved.

In order to settle above-mentioned problems, the present inventors haveaccomplished to design a method for detecting a biomolecule and abiochip therefor, wherein the dielectric constant of a material filledbetween electrodes within a reaction chamber of a biochip is made large,which does not distort actual change of the capacitance value whenreacting on the surface of a sensing electrode, and thus, reflectsaccurately the biological, biochemical or chemical reaction of abiomolecule in a reaction chamber.

SUMMARY OF INVENTION

The present invention aims to settle the above-mentioned disadvantagethat cannot monitor sensitively a capacitance value changed in anelectrode of a biosensor during actual reaction in prior arts. Theobject of the present invention is to provide a method for detecting thepresence and/or the reaction of a biomolecule by accurately monitoringchanges of electrical property after biological, biochemical or chemicalreaction of a biomolecule in a reaction chamber and a biochip providedtherefor.

Particularly, the object of the present invention is to provide a methodfor detecting changes of electrical property, especially changes of animpedance or capacitance value accurately after biological, biochemicalor chemical reaction of a biomolecule in a reaction chamber, by using asensing electrode installed in the reaction chamber and a biochipprovided for this purpose.

More particularly, the object of the present invention is to provide amethod for detecting a biomolecule and a biochip used for this purpose,wherein the dielectric constant of a material filled between electrodeswithin a reaction chamber of a biochip is made large, which does notdistort actual changes of a capacitance value when reacting on thesurface of a sensing electrode and thus, accurately reflects thebiological, biochemical or chemical reaction of a biomolecule in areaction chamber.

DETAILED DESCRIPTION OF INVENTION

In order to overcome the above-mentioned disadvantage of conventionalmethods, it is necessary to discriminate the capacitance value (C_(d))changed by reacting a receptor immobilized onto an electrode with abiomolecule from the capacitance value (C_(t)) of a buffer solution sothat the capacitance value (C_(d)) should not be passed over by thecapacitance value (C_(t)). For this purpose, 1/C_(t) score needs tobecome smaller and namely, C_(t) score needs to become larger.

Meanwhile, the capacitance value is defined as described in themathematical formula 4. Accordingly, the area of electrode should becomelarger, the interval between electrodes should become narrower, or thedielectric constant of a material filled between electrodes shouldbecome larger in order to increase C_(t) score.

In prior arts, the system for increasing sensing sensitivity, in whichthe width of a sensor electrode is decreased in IDE (interdigitatedelectrode) of a biochip in order to increase C_(t), has adopted asdescribed above. However, the decrease of electrode width may not bepreferred in respect of cost and efficiency, since it requires a highlyadvanced microelectrode formation technique.

Considering such a disadvantage, the present invention has attempted asystem for increasing dielectric constant c(or also referred to asdielectric ratio) of a material filled between electrodes in a reactionchamber of a biochip. This method does not require a highly advancedtechnique of microelectrode formation, as well as not distort actualchanges of a capacitance value when reacting on the surface of a sensingelectrode, and can accurately reflect the biological, biochemical orchemical reaction of a biomolecule in a reaction chamber.

In the meantime, the biochip is filled with a reagent for reacting atarget sample between electrodes of a reaction chamber. Accordingly, itis required to increase the dielectric constant of this reagent. But,this reagent for reacting a target sample is difficult to modulate thedielectric constant due to ions.

Generally, the dielectric constant is defined as the ratio of thecapacity measured when a dielectric substance is filled betweenelectrodes to the capacity measured when nothing is filled betweenelectrodes, and the dielectric ratio of any dielectric substance isalways larger than 1. Among liquid, the dielectric constant of water isknown to be largest and approximately 80 at room temperature. However,in order to test a target sample, water such as distilled water ordeionized water cannot be used for reagent use directly.

Therefore, the present invention provides a method for detecting thepresence and/or the reaction of a biomolecule in a target sample,wherein the reaction of the target sample is performed under knownreaction solution, in which a reference fluid having a high dielectricconstant such as water is filled into a reaction chamber before thereaction to measure an impedance value or a capacitance value, aftercompleting the reaction, the reaction solution is removed and saidreference fluid having a high dielectric constant is refilled to measurean impedance value or a capacitance value, and then the impedance valueor the capacitance value measured before and after the reaction iscompared with each other.

Particularly, the present invention provides a method for detecting abiomolecule electrically, which comprises: (a) providing an apparatusfor electrical detection comprising a reaction chamber receiving atarget sample, and a plurality of sensing electrodes located in thereaction chamber for detecting a biomolecule within the target sample;(b) measuring an impedance value or a capacitance value between thesensing electrodes, after introducing a reference fluid having a highdielectric constant into the reaction chamber; (c) providing a reactionsolution and the target sample into the reaction chamber, after removingthe reference fluid having a high dielectric constant from the reactionchamber; (d) reacting the target sample under the reaction solution inthe reaction chamber; (e) removing the reaction solution from thereaction chamber; (f) measuring an impedance value or a capacitancevalue between the sensing electrodes, after introducing said referencefluid having a high dielectric constant again into the reaction chamber;and (g) comparing the impedance value or the capacitance value measuredin the step (b) with the impedance value or the capacitance valuemeasured in the step (f).

In an embodiment of the present invention, the impedance value or thecapacitance value is measured according to frequency in the step (b) andthe step (f), respectively by a measuring device using alternatingcurrent (AC) electric source. In this case, the impedance value or thecapacitance value measured is illustrated as curves with respect to thefrequency value. In case that the curve of the impedance value or thecapacitance value obtained in the step (b) is discriminated from thecurve of the impedance value or the capacitance value obtained in thestep (f), this result is judged to show the presence of the targetbiomolecule or the reaction of the target biomolecule onto the sensingelectrode within the target sample.

In another embodiment of the present invention, the reference fluidhaving a high dielectric constant is preferable to have approximatelymore than 4 of dielectric constant, more preferably to haveapproximately more than 80 of dielectric constant. The reference fluidhaving a high dielectric constant can be distilled water or deionizedwater, but the present invention is not limited hereto. Any fluid can beapplicable, if it has approximately more than 4 of dielectric constantas described above.

In another embodiment of the present invention, the sensing electrodesare preferably made of gold, chrome, copper or aluminum. Preferably, thesensing electrodes are formed of interdigitated electrode, and eachinterval between the sensing electrodes can be approximately 1 to 20 μmand preferably, approximately 1 to 4 μm.

In another embodiment of the present invention, a plurality of thesensing electrodes can be bound with a receptor capable of interactingwith the biomolecule.

In another embodiment of the present invention, the term “biomolecule”is a material that interacts with the receptor combined on the sensingelectrode and can include nucleic acids comprising one or morenucleotides, proteins comprising one or more peptides, amino acids,glycolipids or small molecule compounds, and preferably an antigen, DNA,RNA or PNA (peptide nucleic acid).

In another embodiment of the present invention, the term “receptor” is amaterial combined on the sensing electrode, which interacts with thebiomolecule and can include nucleic acids comprising one or morenucleotides, proteins comprising one or more peptides, amino acids,glycolipids or small molecule compounds, and preferably an antigen, aprobe or a primer.

In addition, in another embodiment of the present invention, thereceptor has a feature to be an oligonucleotide probe, the biomoleculeis the nucleic acid, and the reaction solution is a solution for nucleicacid hybridization. Further, in another embodiment of the presentinvention, the receptor has a feature to be an oligonucleotide primer,the biomolecule is the nucleic acid, and the reaction solution is asolution for PCR amplification.

In addition, the biochip of the present invention used in the method fordetecting a biomolecule can be manufactured in a form of single plate,but can be in a combined form of two plates. In an embodiment of thepresent invention, this biochip can include a DNA hybridization chip ora PCR reaction chip.

In another embodiment of the present invention, the biochip having acombined form of two plates comprises a first plate that has a pluralityof sensing electrodes bound with a receptor interacting with abiomolecule, and a second plate that forms a space of a reaction chamberwith respect to the first plate by being combined with the first platein a predetermined distance.

The first plate has a plurality of sensing electrodes formed on SiO₂insulation layer on N-type or P-type silicon plate, which performselectrical detection in order to identify whether a biomolecule existsand/or reacts or not. The second plate prevents the sensing electrodesof the first plate from contacting outside environment to form a spaceof a reaction chamber by being combined with the first plate.

The reaction chamber is a space where an impedance value and/or acapacitance value are measured between electrodes after being filledwith a reference fluid having a high dielectric constant. In addition,the reaction chamber is a space where a biomolecule in the target sampleis reacted after being filled with the reaction solution and the targetsample. In case that the receptor is immobilized on the sensingelectrode, the biomolecule within the reaction chamber reacts with thereceptor and as a result, the changes of an impedance value and/or acapacitance value between electrodes occur.

The above-mentioned plates can have various sizes, depending upon thekind and the number of applicable receptors, for example the kind andthe number of probes and primers. In addition, the plates can havesquare shape, rectangular shape, and circular shape, but are not limitedhereto in the present invention. Preferably, the first plate may be asilicon plate, and the second plate may be a glass plate. But, theseplates can be also constructed by any one selected from the groupconsisting of glass, silicon, fused silica, polystylene,polymethylacrylate, polycarbonate, gold, silver, copper and platinum.

In addition, the second plate forming the space of the reaction chamberis perforated to have a fluid inlet in the direction of thickness andfurther perforated to have a fluid outlet in the direction of thicknesson the opposite side of the fluid inlet. The inlet and the reactionchamber, and the outlet and the reaction chamber are communicated witheach other to allow fluid to flow therethrough in the reaction chamber.The reference fluid having a high dielectric constant, the reactionsolution and the target sample are introduced through the inlet, and thereference fluid having a high dielectric constant, the reaction solutionand the target sample flow outside through the outlet.

In addition, an inlet micro channel that directs to introduce fluid intothe inlet, and an outlet micro channel that directs to discharge fluidfrom the outlet can be made by layering a third plate on the secondplate. The inlet micro channel is communicated with an injection openinginto which the reference fluid having a high dielectric constant, thereaction solution and the target sample are injected, and penetratesthrough the third plate. The outlet micro channel is communicated with adischarge opening into which the reference fluid having a highdielectric constant, the reaction solution and the target sample aredischarged, and penetrates through the third plate.

In addition, a valve structure that can control fluid injection into thespace of the reaction chamber and fluid discharge from the space of thereaction chamber can be provided on the inlet micro channel and/or theoutlet micro channel. The valve structure is a passage extended afterbeing penetrated into the third plate in the direction of thickness,between the injection opening and the space of the reaction chamber, orbetween the discharge opening and the space of the reaction chamberwherein oil is stored in said passage. When nitrogen gas (N₂) isinjected into this passage of the valve structure, the oil stored in thepassage expands and blocks the channels due to the oil expansion so asto obstruct inflow of fluid from the injection opening to the reactionchamber and outflow of fluid from the reaction chamber to the dischargeopening. Preferably, the third plate is made of PDMS(polydimethylsiloxane).

The micro channel and the valve structure mentioned above can beprepared in a micron unit or nano unit by using a MEMS(micro-electro-mechanical-system) technique used widely in this field,and their size can be controlled, depending upon receptor size,electrode size, or reaction condition.

The channel structure and the valve structure facilitate following stepsin the above-mentioned method: filling the reference fluid having a highdielectric constant into the reaction chamber before reaction; removingthe reference fluid having a high dielectric constant from the reactionchamber and filling the reaction solution and the target sample into thereaction chamber during reaction; and removing the reaction solutionfrom the reaction chamber and refilling the reference fluid having ahigh dielectric constant into the reaction chamber after the reaction.For this purpose, the biochip can be provided with an injection pump anda discharge pump in order to inject and remove the fluid, the reactionsolution and the target sample. For driving a pump, the biochip can alsobe provided with a small motor or a gearbox.

On the other hand, an impedance analyzer can be connected electricallyto a plurality of sensing electrodes. When this biochip is a PCRreaction chip, a PCR device for conducting PCR reaction can bemanufactured to accommodate a biochip and thus, the biochip and the PCRdevice can be provided as one set. Preferably, when the biochip and thePCR device are provided as one set, the PCR device having an impedancemeasuring apparatus is used and the impedance measuring apparatus isconnected electrically to sensing electrodes of the biochip.

ADVANTAGEOUS EFFECTS

As illustrated and confirmed above, the method of present invention candetect the change of electrical property accurately according tobiological, biochemical or chemical reactions of the biomolecule withinthe reaction chamber, even though the capacitance value (C_(d)) changedon the surface of sensing electrodes may be passed over by thecapacitance value (C_(t)) of the buffer solution.

Particularly, the present invention is advantageous to make thedielectric constant of a material filled between electrodes within thereaction chamber of the biochip large, which does not distort actualchange of a capacitance value when reacting the biomolecule in thetarget sample with the receptors immobilized on the surface of sensingelectrodes, and can reflect total change of the capacitance valueaccurately and thus, detect the presence and/or the reaction of thebiomolecule sensitively.

BRIEF DESCRIPTION OF DRAWINGS

The above and other objects, features and other advantages of thepresent invention will be more clearly understood from the followingdetailed description taken in conjunction with the accompanyingdrawings, in which;

FIG. 1 depicts the change of the impedance value measured by using aconventional system for detecting a PCR reaction.

FIG. 2 is a flow chart of process that illustrates a method forpreparing a DNA hybridization chip used in the present invention.

FIG. 3 is a top plane view illustrating the electrode portion of the DNAhybridization chip manufactured according to the process of FIG. 2.

FIG. 3( a) shows the whole electrode portion of the DNA hybridizationchip, FIG. 3( b) depicts schematically a magnified part of the electrodeof FIG. 3( a), and FIG. 3( c) depicts a pair of electrodes (20 a, 20 b)having a finger shape which are partially magnified.

FIG. 4 is a cross sectional view of the metal electrode (20) of FIG. 3(b) sectioned according to A-A′ line.

FIG. 5 is a flow chart of process that illustrates a method forpreparing a PCR reaction chip used in the present invention. FIG. 5(A)illustrates a process for preparing a silicon plate of a PCR reactionchip, FIG. 5(B) illustrates a process for preparing a glass plate of aPCR reaction chip, and FIG. 5(C) depicts the binding state of thesilicon plate and the glass plate.

FIG. 6 is a sectional view that illustrates a PCR reaction chip to whicha fluid controlling plate based upon PDMS is attached. FIG. 6. (A)depicts a fluid controlling plate attached on the glass plate of the PCRreaction chip manufactured previously. FIG. 6. (B) and (C) illustrate aprocedure for controlling fluid flow in the PCR reaction chip by a valvefunction of the fluid controlling plate.

FIG. 7 depicts an electrode array type chip.

FIG. 8 depicts the change curve of the impedance value measured by usingthe method for detecting DNA hybridization reaction in the presentinvention.

FIG. 9 depicts the change curve of the impedance value measured by usingthe method for detecting PCR reaction in the present invention.

EXAMPLES

Practical and presently preferred embodiments of the present inventionare illustrated more clearly as shown in the following examples.

However, it should be appreciated that those skilled in the art, onconsideration of this disclosure, may make modifications andimprovements within the spirit and scope of the present invention.References cited in the specification are incorporated into the presentinvention.

Example 1 Preparation of a DNA Hybridization Chip

In the method for detecting a biomolecule of the present invention, whena target biomolecule is DNA and a receptor combining with a sensingelectrode is an oligonucleotide probe, a DNA hybridization chip forhybridizing a probe and DNA as a biomolecule is prepared.

The DNA hybridization chip sensor is manufactured by using a generalprocedure disclosed in this field, including thin film formationtechnique of silicon dioxide, photolithography technique, exposurepatterning technique, developing technique, wet and/or dry etchingtechnique, etc. Particularly, sensing metal electrodes having an IDE(interdigitated electrode) are formed on a silicon plate by using a MEMS(micro-electro-mechanical-systems) and then, oligonucleotide probes areimmobilized on the sensing electrodes.

The process for preparing the DNA hybridization chip used in the presentinvention will be described according to stages as follows.

Example 1-1 Construction of a Silicon Plate on Which Electrodes areFormed

Hereinafter, referring to FIG. 2, the procedure for forming a plate (10)and an electrode (20) of the DNA hybridization chip (100) used in thepresent invention will be explained briefly.

N-type silicon wafer (10) having 500 μm of thickness was thermallyoxidized to make about 10,000 Å of a SiO₂ layer (12), an oxidizedinsulation layer (See the Step (a) of FIG. 2). Then, the insulationlayer (12) was coated with a photoresist AZ 5214 (14) (See the Step (b)of FIG. 2). On the photoresist (14), a mask having a pattern of metalelectrodes (not shown) was mounted and exposed with ultra violet light.After that, the silicon wafer (10) was immersed in AZ300MIF developingsolution to be developed and etch-treated (See the Step (c) of FIG. 2).On the etched wafer (10), a thin film (20) of chrome (Cr) (300 Å) andgold (Au) (3000 Å) was deposited (See the Step (d) of FIG. 2). Thedeposition of the thin film (20) of chrome and gold was performed byusing chemical vapor deposition (CVD), vacuum evaporation or sputtering.Then, the photoresist layer (14) remained and a part of the thin film onthe photoresist layer (14) remained were removed by using a knownlift-off process so as to make a metal electrode (20) (See the Step (e)of FIG. 2). In order to form a metal electrode in the format of an IDE(interdigitated electrode), a photoresist AZ4620 (16) was additionallycoated and the electrode (20) was patterned to form a fine IDE by usinga known photolithography procedure in this field (See the Step (f) ofFIG. 2).

The DNA hybridization chip (100) manufactured is illustrated in FIG. 3.The DNA hybridization chip (100) depicted in FIG. 3( a) hasapproximately 2.4 mm of diameter (R) in the electrode (20), but isbetter to have as small diameter as possible in order to manufacture anelectrode array type DNA hybridization chip (See FIG. 7). The area ofthe electrode (20) is approximately 3.6 mm².

In FIG. 3( b), the electrode depicted in FIG. 3( a) is magnifiedpartially and illustrated schematically. In order to be understoodeasily, the shape of the electrode (20) is illustrated with a rectangle,but an actual electrode (20) may be round or elliptic. As illustrated inFIG. 3( b), it is confirmed that the sensing electrodes (20) of the DNAhybridization chip (100) used in the present invention have a pair ofelectrodes (20 a, 20 b) wherein they are arrayed alternately in apredetermined interval with a finger-like shape.

On the other hand, the pair of electrodes (20 a, 20 b) having afinger-like shape is magnified partially as illustrated in FIG. 3( c).The width (t) of each electrode (20 a, 20 b) is approximately 10 μm andthe interval (d) between the electrodes is approximately 4 μm.

Example 1-2 Immobilization of Oligonucleotide Probes on an Electrode

Above all, the electrodes (20) of the DNA hybridization chip (100)prepared in Example 1-1 were washed with oxygen plasma (150 W, 100mtorr) for 10 seconds in order to remove impurities on the electrodes(20) completely. Then, one end of a probe DNA (5′ or 3′) was immobilizedon the electrode (20). For this immobilization, a probe5′-SH—(CH₂)₂₄-GCC ATT CTC ACC GGA TTC AGT CGT C-3′, in which a thiolgroup is attached to a carbon chain (—CH₂—) at the 5′terminus of theprobe oligonucleotide, is prepared and fixed on the surface of the metalelectrode (20) by using a SAM (self-assembled method). In detail, 10pmol/μl of SH-ssDNA probe was reacted under MgCl₂ solution condition for18 hours. As a result, —SH group of the probe reacts on the metalsurface to immobilize the probe DNA (22) on the surface of the metalelectrode (20) (See FIG. 4).

Meanwhile, a conductive polymer can be bound to a probe DNA instead of athiol group, and the conductive polymer bound to DNA can be immobilizedonto the surface of the metal electrode by using a SAM method (selfassembled method). All the conductive polymers known in this field, forexample, especially high-conductive polymer including polyacetylene,poly(p-phenylene) (PPP), polypyrrole (PPy), polyaniline, polythiophene,etc. may be used. The conductive polymer can be bound to the probe DNAby using a known binding method regarding binding between a polymer andDNA.

FIG. 4 is a cross sectional view of the metal electrode (20) of FIG. 3(b) sectioned according to A-A′ line and illustrates the immobilizationof the DNA probes (22) on the metal electrodes (20 a, 20 b) formed onthe silicon plate wherein the oxidized insulation layer (12) intervenesbetween the metal electrodes (20 a, 20 b) and the silicon plate (10).

Example 2 Detection of a Target Nucleotide Sequence Using the DNAHybridization Chip

The DNA hybridization chip (100) of FIG. 4 manufactured in Example 1-2was used to examine whether a target nucleotide sequence (5′-GAC GAC TGAATC CGG TGA GAA TGG-3′) is found within a target sample or not.

Above all, deionized water as a reference solution having a highdielectric constant was introduced into a reaction space betweenelectrodes (20) of the DNA hybridization chip (100) and then, theimpedance value between the electrodes (20) was measured at roomtemperature by applying AC power. The applied magnitude of AC power was100 mV and its frequency range was 1 to 32 MHz when the impedance valuewas measured.

After that, the reference solution was removed from the DNAhybridization chip (100) and then, the target sample containing thetarget nucleotide sequence (10 pmol/μl of ssDNA) and TE buffer wereintroduced into the reaction space of the DNA hybridization chip (100).The resultant was incubated for 18 hours at room temperature dependingupon hybridization condition to hybridize the probe DNA (22) immobilizedonto the electrode (20) of the DNA hybridization chip (100) with ssDNAof the target nucleotide sequence (5′-GAC GAC TGA ATC CGG TGA GAATGG-3′).

Again, the buffer solution was removed from the DNA hybridization chip(100) and deionized water was introduced. Then, the impedance valuebetween the electrodes (20) was measured at room temperature by applyingAC power. The applied magnitude of AC power was 100 mV and its frequencyrange was 1 to 32 MHz when the impedance value was measured.

As described above, the impedance value measured before thehybridization reaction and the impedance value measured after thehybridization reaction are illustrated in FIG. 8. In FIG. 8, contrastingto FIG. 1 in case of a conventional method, the impedance value measuredbefore the hybridization reaction and the impedance value measured afterthe hybridization reaction were observed with two different curves, whenthis experiment is conducted by using a positive target samplecontaining the target nucleotide sequence. As illustrated in FIG. 8, itis confirmed that the impedance value should decrease when the targetDNA exists in the reaction solution during the hybridization reaction.Therefore, it is confirmed that the impedance value will decrease as theconductance value increases due to the hybridization from ss-DNA tods-DNA during the hybridization reaction.

In the method of the present invention, the change of the capacitancevalue (C_(d)) after hybridizing a target ssDNA with a probe immobilizedon a sensing electrode should reflect total change of the capacitancevalue (C_(T)) since it is not passed over by the capacitance value(C_(t)) of a buffer solution. Hence, the method of the present inventioncan permit electrical monitoring to detect the hybridization reactionand the presence of target DNA, because two curves of the changes of theimpedance value are observed before and after the hybridization as shownin FIG. 8.

On the other hand, it is observed that the impedance value measured froma negative target sample without a target nucleotide sequence should besmaller than the impedance value of the positive case as shown in FIG.8.

Example 3 Construction of a PCR Reaction Chip

In the method for detecting a biomolecule of the present invention, whena target biomolecule is DNA and a receptor combining with a sensingelectrode is an oligonucleotide primer, a PCR reaction chip foramplifying DNA as a target biomolecule by using a primer and a PCRreagent is manufactured.

Like the DNA hybridization chip sensor, the PCR reaction chip is alsomanufactured by using a general procedure disclosed in this field. Asilicon plate and an electrode portion can be manufactured by using thinfilm formation technique of silicon dioxide, photolithography technique,exposure patterning technique, developing technique, wet and/or dryetching technique, etc. The sensing metal electrodes having an IDE(interdigitated electrode) are formed on the silicon plate by using aMEMS (micro-electro-mechanical-systems) and then, the oligonucleotideprimers are immobilized on the sensing electrodes. In addition, a glassplate attached to the silicon plate is made by using glass wet etchingand sand blasting technique.

The process for preparing the PCR reaction chip used in the presentinvention will be described according to stages as follows.

Example 3-1 Construction of a Silicon Plate, Electrodes and a ClassPlate

Referring to FIG. 5, the process for manufacturing the PCR reaction chip(200) used in the present invention will be explained as below.

1. Formation of a Silicon Plate and Metal Electrodes

N-type silicon wafer (10) having 500 μm of thickness was thermallyoxidized to make about 5,000 Å of a SiO₂ layer (12), an oxidizedinsulation layer (See the Step (a) of FIG. 5A). Ti/Au (300 Å/3000 Å)metal electrode (20) was patterned by using a conventionalphotolithography process, and a part of the electrode, which will beexcluded from an IDE, was removed to form a fine IDE by dry-etchingusing ion milling (See the Step (b) of FIG. 5A). After that, a SiO₂layer (18) of at least 1 μm of thickness was deposited onto thepatterned metal electrode (20) by using chemical vapor deposition (CVD),vacuum evaporation or sputtering, etc. (See the Step (c) of FIG. 5A).Then, the SiO₂ layer (18) of the insulation layer was made even by usingchemical mechanical polishing (CMP) (See the Step (d) of FIG. 5A). Afterthat, the IDE was only treated with dry etching (RIE) or wet etching(BOE) to expose the surface of Au metal electrode. As a result, aportion of the insulation layer (18) binding to the glass plate (30)described below was formed (See the Step (e) of FIG. 5A).

The pattern of the metal electrode (20) is similar to that of the DNAhybridization chip (100) as shown in FIG. 3( a), and a pair ofelectrodes (20 a, 20 b) is arrayed alternately in a predeterminedinterval with a finger-like shape.

The electrode (20) has approximately 8.6 mm of diameter, but it isbetter to have as small diameter as possible in order to manufacture anelectrode array type PCR reaction chip (See FIG. 7). The width (t) ofeach electrode (20 a, 20 b) is approximately 20 μm and the interval (d)between the electrodes is approximately 20 μm.

2. Preparation of a Class Plate and Construction of a PCR Reaction Chip

As illustrated in FIG. 5B, Cr/Au (300 Å/3000 Å) metal film (32) wasdeposited on the upper and the lower surfaces of a conventional glassplate (Pyrex product) (30) used in a semiconductor process by usingchemical vapor deposition (CVD), vacuum evaporation or sputtering, etc.(See the Step (a) of FIG. 5B). By using a general photolithography, athin metal film (32) formed on the upper portion was patterned andwet-etched to form a protective layer (32) that is a part for protectinga glass plate portion processed later (See the Step (b) of FIG. 5B).Then, the other portion of the glass plate (30) that is not protected bythe protective layer (32) was wet-etched to form a space of a reactionchamber in the PCR reaction chip (See the Step (c) of FIG. 5B). All thethin metal films formed on the upper and the lower surfaces were removedand the bottom surface of the glass plate was treated with a blue tape(any blue tape is available if suitable for patterning) patterning (34)in order to form an inlet/outlet (36 a, 36 b) of the PCR reaction chip(See the Step (d) of FIG. 5B). After that, penetration openingsincluding inlet/outlet openings, etc. were made in the direction ofthickness on the glass plate (30) by performing a sand blast processknown in this field (See the Step (e) of FIG. 5B).

The silicon plate (10) previously manufactured on which the metalelectrodes (20) are formed was bound with the glass plate (30) by usinganodic bonding as illustrated in FIG. 5C. The resulting silicon plate(10) and glass plate (30) was diced to obtain a final PCR reaction chip(200).

The anodic bonding is a method for binding a silicon plate and a glassplate, in which cationic ions (for example, Na⁺ ion) present within aglass plate are allowed to move to the opposite direction with respectto the position where the silicon plate and the glass plate are earthed,to form hydrogen bonds on the surfaces of the silicon plate and theglass plate and to bind the two plates by applying hot heat and highelectric field.

In the PCR reaction chip (200) where the two plates are bound to eachother, a plurality of metal sensing electrodes (20) are formed on thesilicon plate (10) (See FIG. 5 and FIG. 6A), and primers (22) for atarget DNA template are immobilized on the electrodes (See FIG. 6A). Inaddition, this silicon plate (10) is bound to the glass plate (30) in apredetermined distance to form a space for a reaction chamber (46) (SeeFIG. 5 and FIG. 6A).

The space of the reaction chamber (46) is a space where the impedancevalue is measured between the electrodes after being filled with areference solution of deionized water. Also, the space of the reactionchamber (46) is a space that accommodates a target sample and a PCRreaction solution to conduct a PCR reaction, in which the target DNAtemplate present in the target sample binds complementarily to theprimers (22) immobilized on the metal electrodes (20) of the siliconplate (10) to proceed the PCR reaction.

Therefore, the change of the impedance value between the metalelectrodes (20) according to PCR reaction will be detected to monitorwhether the target DNA template is present within the target sample ornot.

The glass plate (30) plays a role to protect the space of the reactionchamber (46) from outside environment. In addition, the glass plate (30)is perforated to have a fluid inlet (36 a) in the direction of thicknessand further perforated to have a fluid outlet (36 b) in the direction ofthickness on the opposite side of the fluid inlet (36 a) (See FIG. 5 andFIG. 6A). As illustrated in FIG. 5 and FIG. 6A, the inlet (36 a) and thespace of the reaction chamber (46), and the outlet (36 b) and the spaceof the reaction chamber (46) are communicated with each other to allowfluid to flow therethrough in the reaction chamber (46). Through theinlet (36 a), the reference solution of deionized water, the PCRreaction solution and the target sample are introduced into the reactionchamber (46), and the reference solution of deionized water, the PCRreaction solution and the target sample are discharged from the reactionchamber (46) through the outlet (36 b).

3. Construction of a Fluid Controlling Plate Based Upon PDMS

In order to control fluid flow within the PCR reaction chip (200), afluid controlling plate (40) based upon PDMS is prepared and attachedonto the glass plate (30) of the PCR reaction chip (200) manufacturedpreviously.

The fluid controlling plate (40) of the PCR reaction chip (200) used inthe present invention is preferable to be prepared with a transparentpolymer in order to allow the inside of the chip (200) to be seen.Particularly, it is more preferable to be manufactured with PDMS(polydimethylsiloxane). Advantageously, PDMS is cheaper than siliconeand easy to be manipulated, flexible and resistant to water when boundto another surfaces. Besides, PDMS does not cause any negative effectupon biomolecules due to its biocompatible property. However, it ispreferably deposited with a reactive coating material or treated withgraft copolymerization. Non-specific adsorption of hydrophobiccomponents of cell such as protein may occur, if not coated. Forexample, when an injection opening (42), channels (44 a, 44 b) and adischarge opening (48) are coated with bovine serum albumin (BSA) in thefluid controlling plate (40), the non-specific adsorption of hydrophobiccomponents can be minimized.

Above all, the fluid controlling plate (40) was designed by using a CAD(computer aided design) program. Then, the resulting design was printedon a photomask and impurities were removed. SU-8 of a negative typephotoresist was spin-coated on the silicon wafer. A template for thefluid controlling plate (40) was manufactured by exposure anddevelopment using a photomask that has patterns of an injection opening(42), a valve structure (52), micro channels (44 a, 44 b) and adischarge opening (48). And then, PDMS of liquid phase was poured intothe SU-8 template, solidified by heating and separated from thetemplate. The resulting PDMS fluid controlling plate (40) was cutaccording to the size of the PCR reaction chip (200) and perforated atthe position, where the injection opening (42) and the discharge opening(48), etc. were formed, in a desired size by using a punch. After that,the fluid controlling plate (40) as manufactured above was oxidized withplasma and then, attached to the corresponding position on the glassplate (30) of the PCR reaction chip (200).

FIG. 6A illustrates a longitudinal sectional view of the finallymanufactured PCR reaction chip (200) to which the fluid controllingplate (40) is attached.

When the fluid controlling plate (40) is adhered onto the glass plate(30), an inlet micro channel (44 a) that directs to introduce fluid intothe inlet (36 a), and an outlet micro channel (44 b) that directs todischarge fluid from the outlet (36 b) are formed. An injection opening(42) that introduces a reference solution of deionized water, a PCRreaction solution and a target sample is formed in the fluid controllingplate (40), by being penetrated into the fluid controlling plate. Theinjection opening (42) can introduce fluid into the reaction chamber(46) since it is communicated with the inlet micro channel (44 a). Inaddition, a discharge opening (48) that discharges a reference solutionof deionized water, a PCR reaction solution and a target sample isformed in the fluid controlling plate (40), by being penetrated into thefluid controlling plate. The discharge opening (48) can discharge fluidfrom the reaction chamber (46) since it is communicated with the outletmicro channel (44 b).

Furthermore, passages (52) extended longitudinally after beingpenetrated into the fluid controlling plate (40) in the direction ofthickness, between the injection opening (42) and the reaction chamber(46), or between the discharge opening (48) and the reaction chamber(46) are formed in the fluid controlling plate (40) (See FIG. 6. (A) and(B)). The passage (52) plays a role of valve for the inlet micro channel(44 a) and/or the outlet micro channel (44 b). Silicone oil (50) isstored in the extended passage (52). As illustrated FIG. 6C, if nitrogengas (N₂) is injected into the passage (52) in which the silicone oil(50) is stored, the oil (50) stored in the passage (52) expands andblocks the channels (44 a, 44 b) due to the oil expansion. Therefore,the fluid controlling plate (40) can block and control inflow of fluidfrom the injection opening (42) to the reaction chamber (46), andoutflow of fluid from the reaction chamber (46) to the discharge opening(48).

Example 3-2 Immobilization of Oligonucleotide Primers on an Electrode

Above all, the electrode (20) of the PCR reaction chip (200)manufactured in Example 3-1 was washed for 30 minutes by using Pirahnasolution (mixture of H₂SO₄ and H₂O₂ in 70:30) to completely removeimpurities on the electrode (20).

After that, one end (5′ or 3′) of the upstream primer and one end (5′ or3′) of the downstream primer were immobilized on the electrode (20).Both the upstream primer and the downstream primer can be immobilized onthe electrode (20), otherwise one end of any primer can be immobilizedon the electrode and one end of the other primer can be mixed in a PCRreaction solution to perform a PCR reaction. For immobilization,5′-SH—(CH₂)₂₄-GCC ATT CTC ACC GGA TTC AGT CGT C-3′ of the upstreamprimer and 5′-SH—(CH₂)₂₄-AGC CGC CGT CCC GTC AAG TCA G-3′ of thedownstream primer, in which a thiol group is attached to a carbon chain(—CH₂—) at the 5′terminus of the primer oligonucleotide, is prepared andfixed on the surface of the metal electrode (20) by using a SAM(self-assembled method). In detail, 10 pmol/μl of SH-ssDNA upstreamand/or downstream primers were reacted for 18 hours in TE buffer. As aresult, the —SH group of the primer is reacted with the metal surface toimmobilize the primer ssDNA (22) on the surface of the metal electrode(20). The immobilization process onto the metal surface performed inthis example is substantially identical with that in example 1-2, asshown in FIG. 4.

Example 4 Analysis of PCR Reaction by Using a PCR Reaction Chip

The PCR reaction chip (200) on which primers are immobilized asmanufactured in Example 3-2 (See FIG. 6) was used to examine whether atarget DNA template is found within a target sample and a PCR reactionis performed or not.

Above all, deionized water of a reference solution having a highdielectric constant was introduced into the reaction chamber (46) of thePCR reaction chip (200) and then, the impedance value between theelectrodes (20) was measured at room temperature by applying AC power.The applied magnitude of AC power was 100 mV and its frequency range was1 to 32 MHz when the impedance value was measured.

After that, the reference solution was removed from the PCR reactionchip (200). Then, Promega PCR Core System II PCR Kit (PCR reactionsolution commercially available in order to perform PCR easily; PromegaCo. Ltd., USA) and DNA template were introduced into the reactionchamber (46) of the PCR reaction chip (200). The PCR reaction wasrepeated with 25 to 35 cycles as recommended by Promega, while adjustingtime and temperature in the PCR reaction during each PCR cycle.

Again, the PCR reaction solution was completely removed from the PCRreaction chip (200) and deionized water was introduced into the reactionchamber. Then, the impedance value between the electrodes (20) wasmeasured at room temperature by applying AC power. The applied magnitudeof AC power was 100 mV and its frequency range was 1 to 32 MHz when theimpedance value was measured.

As a result, the impedance value measured before and after the PCRreaction is illustrated in FIG. 9. In FIG. 9, contrasting to FIG. 1 incase of a conventional method, the impedance value measured before thePCR reaction and the impedance value measured after the PCR reactionwere observed with two different curves, when this experiment isconducted by using a positive target sample containing the target DNAtemplate. As illustrated in FIG. 9, it is confirmed that the impedancevalue should decrease when the target DNA template exists in thereaction solution during the PCR reaction. Therefore, it is confirmedthat the impedance value will decrease as the conductance valueincreases due to the PCR reaction from ss-DNA to ds-DNA during the PCRreaction.

In the method of the present invention, the change of the capacitancevalue (C_(d)) after amplifying a target DNA template with PCR primersimmobilized on a sensing electrode should reflect total change of thecapacitance value (C_(T)) since it is not passed over by the capacitancevalue (C_(t)) of a buffer solution. Hence, the method of the presentinvention can permit electrical monitoring to detect the PCR reactionand the presence of the target DNA template, because two curves of thechanges of the impedance value are observed before and after the PCRreaction, as shown in FIG. 9.

On the other hand, it is observed that the impedance value measured froma negative target sample without a target DNA template should be smallerthan the impedance value of the positive case as shown in FIG. 9.

Those skilled in the art will appreciate that the conceptions andspecific embodiments disclosed in the foregoing description may bereadily utilized as a basis for modifying or designing other embodimentsfor carrying out the same purposes of the present invention.

Those skilled in the art will also appreciate that such equivalentembodiments do not depart from the spirit and scope of the invention asset forth in the appended claims.

We claim:
 1. A biochip for detecting a biomolecule electricallycomprising: a first plate that has a plurality of sensing electrodes; asecond plate that forms a space of a reaction chamber with respect tothe first plate by being combined with the first plate in apredetermined distance; and wherein the second plate forming the spaceof the reaction chamber is perforated to have a fluid inlet in thedirection of thickness and further perforated to have a fluid outlet inthe direction of thickness on the opposite side of the fluid inlet; theinlet and the reaction chamber, and the outlet and the reaction chamberare communicated with each other to allow fluid to flow therethrough inthe reaction chamber; and a reference fluid having a high dielectricconstant, a reaction solution and a target sample are introduced throughthe inlet, and the reference fluid having a high dielectric constant,the reaction solution and the target sample flow outside through theoutlet.
 2. The biochip as claimed in claim 1, wherein an inlet microchannel that directs to introduce fluid into the inlet, and an outletmicro channel that directs to discharge fluid from the outlet are formedby layering a third plate on the second plate; the inlet micro channelis communicated with an injection opening into which the reference fluidhaving a high dielectric constant, the reaction solution and the targetsample are injected and penetrates through the third plate; and theoutlet micro channel is communicated with a discharge opening into whichthe reference fluid having a high dielectric constant, the reactionsolution and the target sample are discharged and penetrates through thethird plate.
 3. The biochip as claimed in claim 2, further comprising avalve structure that can control fluid injection into the space of thereaction chamber or fluid discharge from the space of the reactionchamber on the inlet micro channel or the outlet micro channel,respectively.
 4. The biochip as claimed in claim 3, wherein the valvestructure is a passage extended after being penetrated into the thirdplate in the direction of thickness, between the injection opening andthe space of the reaction chamber, or between the discharge opening andthe space of the reaction chamber wherein oil is stored in said passage.5. The biochip as claimed in claim 1, wherein the third plate is made ofPDMS (polydimethylsiloxane).
 6. The biochip as claimed in claim 1,wherein the biochip performs the following steps: (a) first introducingthe reference fluid having a high dielectric constant into the reactionchamber, and then first measuring an impedance value or a capacitancevalue between the sensing electrodes; (b) removing the reference fluidhaving a high dielectric constant from the reaction chamber from step(a) and introducing the reaction solution and the target sample into thereaction chamber; (c) reacting the target sample under the reactionsolution in the reaction chamber as prepared in step (b); (d) removingthe reaction solution from the reaction chamber after step (c); (e)secondly introducing said reference fluid having a high dielectricconstant again into the reaction chamber, and then secondly measuring animpedance value or a capacitance value between the sensing electrodes;and (f) comparing the impedance value or the capacitance value measuredin the step (a) with the impedance value or the capacitance valuemeasured in the step (e).
 7. The biochip as claimed in claim 1, whereinthe reference fluid having a high dielectric constant has at least 4 ofdielectric constant.
 8. The biochip as claimed in claim 7, wherein thereference fluid having a high dielectric constant is distilled water ordeionized water.
 9. The biochip as claimed in claim 1, wherein thesensing electrodes are formed of interdigitated electrode.
 10. Thebiochip as claimed in claim 9, wherein each interval between the sensingelectrodes is 1 to 4 μm.
 11. The biochip as claimed in claim 1, whereinthe sensing electrode is bound with a receptor interacting with thebiomolecule.
 12. The biochip as claimed in claim 11, wherein thereceptor is a probe or a primer.